Porous tissue scaffolding materials and uses thereof

ABSTRACT

A biodegradable scaffold includes a protein, a cross-linking agent, 10 to 50 percent by volume porosity imparting polymeric beads and a modifying agent. A method for the use of the biodegradable scaffold is also provided.

This application claims the benefit of Provisional Application No.60/122,302, filed Mar. 1, 1999.

GRANT REFERENCE

This research was supported in part by CDC grant R49-CCR411774.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to the fields of biomedical engineeringand the chemistry of wound healing and biological glues. Morespecifically, the present invention relates to a novel porous tissuescaffoldings, which enhance wound healing and also serve as biologicalglues.

2. Description of the Prior Art

Skin injuries are a primary cause of death in North American for peoplebetween the ages of 1 and 44. Although different in etiology andprognosis, the treatment goal for all types of wounds is the same: skinregeneration. Two of the most common skin wounds are skin ulcers andbums. The most prevalent skin ulcers are pressure ulcers, diabeticulcers, and venous statis ulcers.

There are various types of skin wounds. For example, pressure ulcers,caused by prolonged excessive pressure, and burns, caused by excessiveheat, are significant problems in this society. Both heal in theclinical setting, once the causative agent is removed, althoughimproving the speed and quality of healing is highly desirable.

Pressure ulcers are found in 20-30% of the approximately 200,000 spinalcord injury patients, 3-15% of nursing home residents or elderlypatients, and in 3-11% of acute injury patients; for a total ofapproximately 800,000 patients per year. It is estimated that patientswith pressure ulcers incur at least an additional $15,000/year in healthcare costs with estimates as high as $58,000. Enhancing the rate ofregenerative healing would reduce the likelihood and effect of secondarycomplications.

It has been estimated that more than 500,000 persons are treated in ahospital emergency department each year due to thermal injury.Additionally, as many as 70,000 to 100,000 are hospitalized, and ofthese, 10,000 to 12,000 will die. Development of more active treatmentscan accelerate healing in these acute wounds as well as chronicnon-healing; wounds and reduce morbidity and mortality associated withthese skin wounds. Although little is known about the relativeimportance of and the relationships among the factors which enhanceregeneration, it appears that the use of growth factors is one of themost powerful and direct methods presently, available to enhance andcontrol wound healing. A degradable matrix used to deliver the growthfactor, not only protects the growth factor which has a short in vivohalf-life until release, but also serves as a scaffold for tissueformation.

Many investigators have examined ways to enhance and control healing.Healing is affected by altering the wound environment (oxygen, magneticfields, stress, location, etc.) or wound biochemical activity (growthfactors, growth hormone, and other biochemical agents). For skin wounds.rapid regeneration of connective tissue and the overlying epidermis isthe goal. When a wound dressing or implant is used, wound healing can bealtered by changing the implant configuration (pore size, porosity,fiber diameter etc.), the implant surface (composition, charge, surfaceenergy, etc.), the implant biochemical activity (incorporation of growthfactors or other biochemical factors), or the implant physical activity(degradation rate and drug delivery rate). Implants, used in the skin astissue scaffolds, should be degradable to allow complete skinregeneration. When used beneath grafted skin, they should also allowrapid blood vessel infiltration and re-attachment (angiogenesis) betweenthe graft and the underlying, tissue bed. In both cases, theinterconnected porosity is critical for angiogenesis, since bloodvessels require at least 40 mm pores to grow into a biomaterial.Additionally, an adhesive biodegradable matrix would help speed up thesurgical procedure, of skin grafting, while enhancing juxtapositionbetween a skin graft and the underlying tissue bed.

One such biodegradable matrix is fibrin. Fibrin derived from blood hasbeen used as a tissue adhesive. It is generally supplied as atwo-component kit consisting of human-source fibrinogen/Factor XIII andbovine thrombin/CaCl₂. These fibrin sealants have been used since 1972in Europe where a commercial version is presently available. Nocommercial product is available in the United States, and studies havebeen done using autologous or single donor preparation. Clinically, thefibrin matrix has been used as a hemostatic agent, for tissueanastomosis, as a fluid barrier, as a drug delivery vehicle, and as atissue scaffold. Fibrin sealant used for skin grafting has been shown toincrease attachment strength, to the wound bed, compared with staples,leading to less seroma formation and wound contraction.

Another biodegradable matrix suitable for use in wound healing and as abiological scaffold is modified polyethylene glycol (PEG) crosslinkedalbumin.

Using natural biomaterials, such as fibrin and albumin, delivery of abiochemical agent can be accomplished in a number of different ways. Forexample, the fibrin or PEG crosslinked albumin can be impregnated withthe biochemical factor or agent; the agent can be attached to thepolymer chain or it can be included through intra-fibril entrapment.Certain growth factors like FGF can bind to these polymers, like fibrin,Otherwise, the growth factors need to be attached through variouslinkages. When the growth factor is bound, it is released only when thefibrin or albumin degrades. Since the degradation is cellular, thegrowth factor release is controlled by the rate of phagocytic cellularinfiltration and is thus under biofeedback control. Additionally, thedegradation is at the wound edge and thus gives the appropriate gradientto stimulate further angiogenesis and tissue healing.

The use of a degradable matrix to deliver a biochemical agent, such as agrowth factor, protects the growth factor until release, since it seemsto have a short half-life in vivo even in the presence of heparin. Thisshort half-life in vivo is potentially a problem for clinical studiesusing topical administration of growth factors, besides theinconvenience and added personnel expense.

Fibroblast growth factor (FGF), has been shown to induce angiogenesis.The mitogenic, chemotactic and differentiation properties of this growthfactor suggests that it is involved in embryonic development and isprobably a major trophic factor operating at all stages ofembryogenesis. Unlike other growth factors such as platelet derivedgrowth factor (PDGF), transforming growth factor, and epidermal growthfactor (EGF), FGF can stimulate in vivo as well as in vitroproliferation of all cell types involved in wound healing. In, vitrostudies have demonstrated that FGF inhibits contraction while enhancingwound healing. The inhibition of contraction may have therapeuticimplications in the prevention of contracture scars. FGF has also beenshown to increase graft survival by stimulation of epithelialization bycultured keratinocytes and vascularization in the wound bed in athymicmice.

Both basic FGF (FGF-2) and acidic FGF (FGF-1) are extremely angiogenicin vivo and mitogenic for fibroblasts in vitro. A preferential responseby keratinocytes to FGF-1 in either the presence or absence of heparin,compared to FGF-2, has been reported. Stimulation of angiogenesis,granulation tissue formation and neo-epithelialization as well asincreased wound strength, with increased cellularity and collagendeposition, without promoting contraction have been demonstrated inresponse to FGF-1, in vivo, in dermal wounds.

Clinical measures of wound healing include estimates of skin graft take,area healed, time to subjective healing, and length of hospital stay.Even the more quantitative measurement of change in surface area healedsuffers from variability due to the effect of wound size on healingrate. The larger the wound, the more tissue that can be laid down ineach time period, and the larger the change in volume or surface area.The optimal parameter to assess rate of healing should be independent ofwound size. A measure that fits this description is the change inaverage wound diameter. It has been shown that the epidermal migrationrate is relatively constant at approximately 1-2 mm/week.

Additionally, skin grafts secured with sutures or staples are commonlyused in the treatment of patients with third degree bums. The associatedpotential for fluid collection between the wound surface and the graftedskin, which comprises the viability of the graft, may be minimized byattaching grafts using naturally occurring glues. Also, these gluesprovide a scaffold for tissue and can be impregnated with growth factorsand antibiotics to enhance wound healing. In addition, the rate ofdegradation can often be controlled by biofeedback. Fibrin glue is apopular choice for applications of this nature and has recently beenapproved for hemostatic uses in the U.S. The current cost ($300/2 ml)has precluded the widespread use of this product in the U.S. Hence,applicants are examining other commercially viable alternatives. Onesuch possibility is albumin glue, which is FDA approved, does not havethe disease concerns associated with fibrin, and is considerablycheaper.

The prior art is deficient in the lack of effective means offacilitating the repair or regeneration of biological structures such asin the healing of wounds. The present invention fulfills thislongstanding need and desire in the art.

SUMMARY OF THE PRESENT INVENTION

Therefore, it is an object of the present invention to designbiodegradable scaffold or porous materials which are capable of locallydelivering at least one agent including fibroblast growth factor(FGF-1). The present invention also provides a bioadhesive scaffold orporous material including albumin, a polyethylene crosslinking agent,and a modifying agent incorporated into the bioadhesive material.

Other and further aspects, features, and advantages of the presentinvention will be apparent from the following description of thepresently preferred embodiments of the invention given for the purposeof disclosure.

BRIEF DESCRIPTION OF THE DRAWINGS

A better understanding of the present invention will be had uponreference to the following detailed description when read in conjunctionwith the accompanying drawing, and in which:

FIG. 1 illustrates the shear strength of PEG crosslinked albumin andfibrin;

FIG. 2 is a histogram illustrating the percent of wound area filled withremaining material wherein bars represent standard mean error (SEM);

FIG. 3 is a histogram illustrating the percent wound filled with newtissue wherein bars represent standard mean error (SEM);

FIG. 4A is a histogram illustrating the percent inflammatory cells (ofall cells in the wound region) by week wherein bars represent standardmean error (SEM);

FIG. 4B is a histogram illustrating the percent inflammatory cells bymaterial of all cells in the wound region wherein bars representstandard mean error (SEM);

FIG. 5 is a diagrammatical illustration of the test set-up forevaluating shear strength of various adhesives;

FIG. 6A is a force-deformation plot wherein the area under the plot wasconsidered for calculation of failure energy (W) from force-deformationcurve;

FIG. 6B is a plot wherein the shear modulus (G) is calculated from astress-strain curve;

FIG. 7A illustrates polyethylene glycol particles with an averagediameter ranging between 10-120 μm, MW=100,000;

FIG. 7B illustrates polyethylene glycol particles having an averagediameter between 120-200 μm, MW=100,000;

FIG. 8A illustrates the resultant clots formed when particles are addedto adhesive with 40-120 μm average diameter particles added;

FIG. 8B illustrates the resultant clots formed when particles added tothe adhesive having 120-200 μm average diameter particles were added;

FIG. 9 is a graph illustrating a stress-strain curve; and

FIG. 10 is a graph illustrating the mean shear strain at ultimatestrength wherein bars represent standard mean error (SEM) and *indicates significance (P<0.05) with respect to non-porous material.

DETAILED DESCRIPTION OF THE PRESENT INVENTION

In accordance with the present invention there is provided bioadhesivescaffold materials, which are biodegradable and serve as porousdegradable tissue scaffolds and drug delivery mechanisms for the repairand/or regeneration of biological structures. The bioadhesive scaffoldsin accordance with the present invention can include fibrin andmodified-polyethylene glycol (PEG) crosslinked albumin.

These porous degradable tissue scaffold materials can be utilized ashemostatic agents, tissue adhesives, drug delivery matrices, tissuescaffolds, and combinations thereof in the areas of neurology (toreconnect nerves), orthopedics (for attachment of cartilage, bone, andimplants), dentistry (for bone and implants), burns and trauma (for skingrafts), pressure ulcers, ophthalmology (corneal overlays), generalsurgery (replacement of surgical sutures), gastroenterology, and thefilling of a fistula or defect in general, dental, and orthopedicsurgery. The bioadhesive material can be utilized as a general tissuesealant to prevent leaks or as a bonding agent.

The present invention provides a bioadhesive scaffold composition whichhas high mechanical strength, flexibility, fast cure rate, andsufficient adhesion to bond and/or seal tissue in vivo.

The preferred scaffold materials for the present invention include serumalbumin protein and fibrin.

The crosslinking agents can include any suitable crosslinking agentknown to those skilled in the art. However, the preferred crosslinkingagent for the present invention is polyethylene glycol (PEG) or modifiedpolyethylene glycol such as polyethylene glycol succinimidyl propionate(PEG-SPA). The polyethylene glycol can also be used in bead form to makethe matrix porous.

The components which make up the bioadhesive scaffold materials arepreferably mixed together and immediately applied and/or used in thedesired application. The fibrin is preferably present in a concentrationof 1-60 mg/ml final concentration. The albumin is preferably present inan amount ranging from approximately 0.1% to 50% by concentration(0.01-0.50 g/ml). More preferably, the albumin is present in an amountof approximately 1.0% to 30% concentration and, most preferably, thealbumin is present in an amount ranging from 10-30% concentration. ThePEG can be used for three separate structural changes to the albumin orfibrin: to attach factors to the polymer (fibrinogen or albumin), tocrosslink the polymer, or to be used as water soluble beads to createporosity. The PEG to attach factors can be used to get one to sixattachment sites per molecule. For fibrinogen, one attachment permolecule is preferred. The PEG crosslinker is present in an amountranging from approximately 0.1% to 30% by concentration. Preferably, thePEG is present in an amount ranging from between approximately 5% and25% concentration. The PEG for imparting porosity uses 10-50% by volumeof 10-500 μm (micron) PEG beads (5,000-2,000,000 MW). Preferably, thePEG beads are in the 10-20% by volume of 120-200 μm beads(20,000-100,000 MW). It should be noted, that the relative amounts offibrinogen, albumin, crosslinking agent (PEG), and PEG beads may bevaried in order to alter the properties of the bioadhesive scaffoldmaterial including the curing time and the viscosity. That is, therelative amounts of the components may be altered, for example, in orderto tailor the bioadhesive material to a specific application or methodof applying the bioadhesive material.

The bioadhesive scaffold materials can be applied by any suitabletechnique including the use of an applicator, injected via a syringe orsprayed.

As briefly discussed above, the bioactivity of the bioadhesive scaffoldmaterials can be altered by the incorporation of factors or agents, suchas pharmaceuticals, into the bioadhesive scaffold material which aresubsequently released over time as the bioadhesive material degrades.Exemplary agents or factors include growth factors such as FibroblastGrowth Factor I (FGF-1), anti-inflammatory agents, antibiotic agents,and pore forming agents.

The following examples are given for the purpose of illustrating variousembodiments of the invention and are not meant to limit the presentinvention in any fashion.

EXAMPLE 1 Manufacture of Porous Material

Pooled donor fibrinogen and bovine thrombin was obtained from theAmerican Red Cross blood bank. The fibrinogen cryoprecipitate wasre-suspended at 37° C. with 3.3 ml of ddH₂O to obtain a concentration of120 mg/ml. The thrombin solution was prepared at 1000 units/ml in a 80mM CaCl₂ dihydrate solution. The fibrinogen was mixed with thrombin in a1:1 ratio, giving an ultimate fibrinogen concentration of 60 mg/ml.

Polyethylene glycol (PEG)(MW=20,000 Daltons) beads of 100-22 μm diameter(Shearwater Polymers, Inc. Huntsville, Ala.) were added to the thrombin.The 1.2 ml of PEG corresponded to a total of 12% by volume addition tothe final fibrin clot. This volume was chosen because studies haveindicated that 12% volume of the beads ensures interconnecting porosity,(Mikos et al., 1993).

EXAMPLE 2 In Vivo Testing

Five 6 mm diameter ulcers up to the depth of bare cartilage were createdin each rabbit ear after routine surgical preparation, using a trephine.There were four rabbits in the study. There were two periods of study(four and eight days), with two rabbits in each study period. There werefifteen ulcers in each time period, five of which were treated with aporous fibrin scaffold, five with a non-porous fibrin scaffold and fivewere untreated to serve as controls.

All the ulcers were handled using sterile procedures to preventinfection. After the rabbits recovered from surgery, they were housedunder veterinary supervision and returned to individual cages where theywere checked and fed daily. Clinical evaluations were made daily and theulcers were monitored. At the predetermined day, the rabbits from eachtime period were sacrificed by barbiturate overdose. The ulcer andsurrounding tissues were removed. The specimens were then placed in anacid alcohol fixative for light microscope studies.

The implants were embedded in paraffin, sectioned, and stained.Hematoxyline and eosin staining was used for routine observation ofcells and Masson's trichrome stain for measuring collagen formation. Thestained tissue sections were evaluated by histomorphometrical methods toobtain the volume fractions of the various cellular and tissuecomponents. This technique allows volume and other 3-D parameters to beobtained from 2-D measurement of the object. The volume percentage ofcell types {macrophage (Mv), fibroblasts (Fv) and polymorphonuclearleukoctyes (Nv)}, tissue composition {apparent blood vessels (Bv) andcollagen (Cv)}, and number of blood vessels (NB) per field weredetermined at the ulcer edge and the center of the ulcer. Mv, Nv, and Fvwere determined using photographs of fields taken from the lightmicroscope. Bv, NB and Cv were determined using an image analysis system(IBAS, Kontron Image Analysis Division, Munich, Germany). Because ofpossible shrinkage of blood vessels during the histological processing,the volume fraction of bloods vessels were categorized as “apparentvolume fraction.”

Data from four random fields (from at least two different sections) foreach ulcer, both center and edge, were average to give one value perparameter for each ulcer. The epithelialization rate (ER) and thecontraction rate (CR) of the healing ulcer was determined.Epithelialization rate was calculated by measuring the length of newepithelial tissue and dividing this value by the number of weeks in thetime period (4/7 or 8/7), thus giving the value in mm/week. Newepithelium is epidermal cells growing over the ulcer, which can bedistinguished by the type of collagen and presence of hair follicles.Contraction rate was determined by measuring the original ulcer size,and subtracting that value from the diameter of the initial ulcer size(6 mm). The resulting value was also normalized by dividing by number ofdays in the time period, thus giving the value in mm/week. These datawere then used to compare the three treatment groups. The criticalfactors were the extent of angiogenesis (NB, Bv), the fibroblastic (Fv)response and epithelial coverage. SAS version 5 (SAS Institute, Cary,N.C.) software program was used to perform analysis of variance andleast square means, using the general linear model, to detectsignificant differences among the 6 groups. All tests were two tailed,with a level of significance of p <0.05.

Although the histomorphometric data at day four showed differences amongthe groups, the only statistically significant difference was betweenthe control and the treatment groups was in Cv. The Cv for the porousfibrin treated group was about 80% higher than that of the control groupat both the center and the edge of the ulcer. There were nostatistically significant differences between the edge and the center ofthe ulcer for any of the parameters at this time period.

Several histomorphometric parameters, Nv, Mv, Fv, Cv and NB, showedsignificant differences between the treatment groups and the control onday eight. The Nv for both the center and edge of the ulcer in thenon-porous treated group was highest among all three groups, but bothtreated groups were higher than the controls. The Nv for the controlgroup was about 80% lower, and the Nv of the porous fibrin treated groupwas about 60% lower than that of the non-porous fibrin treated group atthe center of the ulcer. At the edge of the ulcer, the Nv for thecontrol group was about 82% lower, the Nv of the porous fibrin treatedgroup was 73% lower than that of the non-porous fibrin treated ulcers.There were also significant differences between the center and edge ofthe ulcer in the porous fibrin treated groups. The Nv at the edge of theulcer was about 51% higher than that seen at the center of the ulcer.

The Mv at the center of the ulcer for the control group was about 70%lower than that seen for the non porous scaffold treated ulcers. The Fvfor the control ulcers at both the center and the edge of the ulcers wasabout 88% lower than that seen for the fibrin treated groups. The Cv forthe control ulcers was about 110% higher than in the treatment groups atboth the center and the edge of the ulcers. The NB for the porous fibrintreated ulcers was significantly higher at the center and at the edge ofthe ulcers than both the non-porous and the control groups. The NB perfield for the porous scaffold was about 400% higher than control and225% higher than the non-porous scaffold.

Between four and eight days, the fibrin treated groups showed asignificant decrease in Cv and an increase in Nv and Fv. The controlgroup, however, showed a significant increase in Cv at the end of dayeight. In the porous fibrin treated group, the NB at the termination ofthe study was significantly higher than that seen at four days. The Mvat the edge of the ulcer also showed a significant increase at the endof the study. Epithelialization rate in all the groups was significantlyhigher at eight days compared to that at four days. Overall there were afew statistically significant differences in histomorphometricmeasurements between the ulcer edges and the ulcer center. Thesedifferences were seen only after eight days.

The rabbit ear ulcer model permits the study of wound healing withoutthe confounding effects of contraction (Mustoe et al., 1991) and higherskin mobility normally seen in dorsal rabbit skin models (Pandit et al.,1994). The rabbit ear is an excellent model to study angiogenesisbecause it is relatively avascular.

The n value, which determines the sensitivity of the study, wascarefully considered in the experimental design. For example, with the nvalue of 5, as used herein, the detectable difference for each parameteris approximately 1.5 times the standard deviation. With standarddeviations averaging about 50% of the mean, changes in the mean of about75% would be needed to detect significant differences. This differenceis relevant to clinical significance. Although this design will notdetect small differences in healing, it was selected on the basis ofclinical significance. For example, the healing times for full-thicknessburns and pressure ulcers would have to be significantly reduced to makea new treatment preferable over the standard skin grafting techniques.In the case of pressure ulcers, if healing rate could be cut in half, itwould approximate the six weeks of hospital bed rest that normallyfollows skin graft surgery, rather than the usual three to six monthsfor conservative treatment.

Little contraction was seen; thus, overall healing can be approximatedby the percent of the ulcer covered by epidermal cells. Al the end offour days, the ulcers were about 20-30% epithelialized, and 90-95%epithelialized at the end of eight days.

The histomorphometric parameters used were indicators of the healingresponse. To decide upon the optimal system to be used in clinicaltrials, it was necessary to prioritize the individual histomorphometricparameters. The overall healing rate, approximated by theepithelialization rate, is the most important clinical measure.Therefore, this parameter was given the highest priority. Because of theimportance of angiogenesis (Bv, NB) and fibroblasts (Fv and Cv) in thescaffolding effect of the matrix, these parameters were given the nextpriority. The time course of the inflammatory response (Mv and Nv) isalso critical in determination of the optimal system.

This wound healing model has an extremely low contraction rate. Incomparison, in a full-thickness dorsal defect model, wound closureduring the first week was predominantly by contraction (CR—5.0 mm/wk) asopposed to epithelialization (ER—1.2 mm/wk), giving a CR/ER ratio ofabout 4:1. In the present invention, this ratio was approximatelyreversed with a 1:6 ratio at day four and a 1:4 at day eight for thecontrol ulcers.

As increase in NB at both the ulcer center and edge by day eightindicated a higher angiogenic response in the porous scaffold treatedulcers. NB in three dimensions, based on histomorphometric principles,gives length per unit volume of blood vessels. Although notstatistically different (p=0.06), the Bv for the porous scaffold at dayeight was about 140% higher than the control at the center of the ulcer.The relatively high standard deviation for this measure (about 75% ofthe mean of the porous scaffold) made a difference in means of thismagnitude too low to be detected in this design. The use of a higher nvalue or a one-tailed test in the original design would have made thisdifference statistically significant. No significant differences inangiogenic response was observed at day four. The enhancing effect ofthe scaffold appears to occur after the ulcer is about 30% healed.

A higher fibroblastic (Fv) reaction, as seen in both the fibrin scaffoldtreated groups, gave an indication of the amount and rate of collagenthat could be produced and the quality of the ulcer. The Cv, however,which indicates collagen density, was higher for the porous system atday four but highest for the control at day eight. Both collagen densityand healed ulcer volume would be needed to determine the total amount ofcollagen produced. The higher Nv and Mv reaction, when healing wasnearly complete, indicated an enhanced inflammatory reaction for thefibrin scaffolds.

These results indicate that the healing response was modified by afibrin scaffold. Both porous and non-porous fibrin scaffolds initiated ahigher fibroblastic response, with the porous fibrin scaffoldsinitiating the highest angiogenic and fibroblastic response. A higherinflammatory reaction was also seen in the non-porous fibrin scaffoldcompared to the porous scaffold.

That a 100-200 μm degradable porous system leads to a significantincrease in the angiogenic and fibroblastic response with acorresponding decrease in healing time as compared to untreated ulcersor non-porous systems was partially shown. The increase in angiogenesisand Fv did not correspond to an increase in epithelialization rate.Modifications to the system to increase the epithelialization rate wouldinclude optimization of the implant's scaffolding ability, both itsconfiguration and bioactivity. This would include optimization of theimplant porosity or other aspects of configuration to achievesignificant enhancement of regenerative skin healing. Addition ofbiochemical factors (e.g., growth factors), alteration in the woundenvironment by changing the oxygen gradient and electromagnetic fieldsalso are important in the modulation of epithelialization rate.

EXAMPLE 3

In Vivo Dose Response of Acidic Fibroblast Growth Factor DeliveredThrough a Porous Fibrin Scaffold on Wound Healing

In order to enhance the scaffolding of a biomaterial both thebioactivity and configuration of the implant was examined. Parameters ofbioactivity include rate of implant degradation as well as the use ofincorporated biochemical factors such as growth factors. Parameters ofimplant configuration includes both surface roughness and porosity. Aporous fibrin scaffold with a pore size of 100-200 μm initiated anenhanced fibroblastic and angiogenic response over the control ulcers.

EXAMPLE 4 Growth Factor Manufacture and In Vivo Dose Response

FGF-1 was expressed in Escherichia coli under the control of trp-lacpromoter. Synthetic DNA fragments encoding the entire frame of humanFibroblast Growth Factor-1 and its NH₂-terminal truncated form wereconstructed and then expressed in E. coli.

The growth factor was added to fibrinogen prior to mixing, at theappropriate concentrations. The amount of FGF-1 added was equal to theamount desired to be released over a one week interval. For example, ifit was desired to release 10 mg/ml PGF-1 to the ulcer each day for sevendays; 70 mg (10×7) of FGF-1 was incorporated into the fibrin clot thesize of the ulcer (6 mm or approximately 28.3 mm²) Based on preliminarystudies, approximately 0.14 ml would produce a clot to cover this size.Therefore, the concentration would be 2.47 mg/mm² (70 mg/28.3 mm²) or500 mg/ml (70 mg/0.14 ml) of fibrongen solution.

EXAMPLE 5 In Vivo Testing of Growth Factors

Five 6 mm diameter ulcers up to the depth of bare cartilage using atrephine were created on each rabbit ear after routine surgicalpreparation. A control ulcer with no growth factor, treatment along withfour different combinations of the growth were tested. The growth factorwas added in the following concentrations: 0.8, 8, 80 and 800 mg/ml.There were six rabbits in the study in this two period study (four andeight days) with three rabbits in each period of study. In each earthere were four ulcers and a control ulcer. Hence, there were five ofeach type of treatment per time period. Each ulcer was covered by aconventional “Opsite” (Smith & Nephew, Inc., Massillon, Ohio) wounddressing. Since it is expected that the growth factor will acceleratethe growth rate, four and eight day tests were performed, less than theten days required for complete healing of control wounds in earlierstudies. The ulcers were handled using sterile procedures as describedabove.

EXAMPLE 6 Necropsy

At the predetermined termination day, the rabbits from each time periodwere sacrificed using barbiturates. The ulcers were removed and wereplaced in an acid alcohol fixative for light microscope studies. Theimplants were embedded in paraffin, sectioned, and then stained.Hematoxylin and eosin, for routine observation of cells, and Masson'strichrome, for collagen were used. To obtain the volume fractions of thevarious cellular and tissue components, the stained tissue sections werethen evaluated using histomorphometry. This technique allows volume andother 3-D parameters to be obtained from 2-D measurements of the object.The volume percent of cell types {macrophage, fibroblasts andneutrophils} and tissue composition {volume fraction of blood vessels}were determined at the wound edge and the center of the wound. Inaddition, the number of blood vessels, epithelialization rate and thecontraction rate of the healing ulcer was quantified. Mv, Nv, and Fvwere determined using photographs of fields taken from a lightmicroscope. Bv and NB were determined using an image analysis system(BAS, Kontron Image Analysis Division, Munich, Germany). Because ofpossible shrinkage of blood vessels were categorized as “apparent volumefraction”. Data from four random fields (from at least two differentsections) for each wound were averaged to give one value per parameterfor each wound. In addition, epithelialization rate and contraction rateof the healing ulcer were quantified.

Epithelialization rate was determined by measuring the length of newepithelial tissue and dividing this value by the number of weeks in thetime period (4/7 or 8/7); thus giving the value in mm/week. Newepithelium is epidermal cells growing over the wound which can bedistinguished by the type of collagen and presence of hair follicles.Contraction rate was determined by measuring the original wound size;and subtracting that value from 6 mm (the diameter of the initial ulcersize). The resulting value was normalized by dividing it by number ofdays in the time period; thus giving the value in mm/week. This data wasanalyzed to determine the best FGF1 concentration. The critical factorswere the extent of angiogenesis (NB, Bv), the fibroblastic (Fv) responseand epithelial (epithelialization rate).

EXAMPLE 7 Volume Fraction of Neutrophils (Nv)

At day four, the Nv at the center of the wound was significantly lowerin the control wounds than all the treatment groups. Nv at the edge ofthe wound was, however, also lower than all the treatment groups exceptfor 0.8 mg dose. There were no significant differences seen after the8th day among any of the treatment groups. The Nv on day eight in thecontrol group was about 200% higher than that seen on day four. Althoughthere were differences seen between the two time periods in the othertreatment groups, none of these were statistically significant.

EXAMPLE 8 Volume Fraction of Macrophages (Mv)

Mv, at the edge of the wound, in the control group after the four daytreatment period was significantly higher than the treatment groups.Similar results were observed at the center of the wound with theexception of Mv for the 80 mg dose. At eight days no significantdifferences were observed between the treated groups and the controls.However, Mv for the 0.8 mg dose was about 110% higher than that seen forthe 800 mg dose. The Mv on day four in the control group was about 75%higher than that seen on day eight. Although there were differences seenbetween the two time periods in the other treatment groups, one of thesewere statistically significant.

EXAMPLE 9 Volume Fraction of Fibroblasts (Fv)

For FV, there were no significant differences seen at the center of thewound at the end of four days, except in 8 mg where the Fv was about 60%lower than the control. The Fv at the center of the wound in the 8 mgtreated groups however, was significantly higher than all the groups atthe end of the eight day period. On day eight, 8 mg dose showed about a310% increase in Fv at the center of the ulcer than that seen on dayfour and a 70% increase compared to the edge of the ulcer on day eight.Although there were differences seen between the two time periods in theother treatment groups, none of these were statistically significant.

EXAMPLE 10 Volume Fraction of Blood Vessels (Bv)

With respect to Bv, the control wound showed a lower response comparedto the all the growth factor-treated groups on day four. The onlysignificant differences were seen at the 8 mg and 0.8 mg dose. Among thegrowth factor treated wounds, the 8 mg treated defects showed asignificantly higher Bv compared to all the other doses with theexception of 0.8 mg. There was no significant difference between thesetwo doses although higher volume fraction of blood vessels were seen inthe 8 mg treated defects. On day four, the 8 mg dose showed about a 200%increase in Bv at the center of the ulcer than as on day eight and a120% increase than at the edge of the ulcer on day four. Also on dayfour, a 0.8 mg dose the Bv at the center of the wound showed a 160%increase than that seen on day eight.

EXAMPLE 11 Number of Blood Vessels (NB)

Significant differences were seen in the NB between the two time periodsin the control ulcer at the edge of the wound. NB at the edge of thecontrol wound at day eight was about 360% higher than that seen on dayfour. NB in the control group at the center of the wound was lower thanall the growth factor doses. However, only the 80 mg and the 0.8 mg dosewere significantly higher than the control. Similar results were seen atthe edge of the wound with 0.8 mg dose being significantly higher thanthe control. Although there were differences seen between the two timeperiods in the other treatment groups, none of these were statisticallysignificant.

EXAMPLE 12 Epithelialization Rate (ER)

Epithelialization rate in all the growth factor treated ulcers wassignificantly higher than that seen in the controls at the end of dayfour. The epithelialization rate in the FGF-1 treated ulcers were about200% higher than that seen in the control. Epithelialization rate at theeight day time period in all the treatment groups with the exception of80 mg and 0.8 mg were significantly higher than that seen on day four.No other significant differences were seen in the contraction rate amongthe treatment groups.

The present invention demonstrates the effective dosing regimen of FGF-1through a porous fibrin carrier vehicle. Histomorphometric parameterswere used as an indication of the healing response. It was necessary toprioritize the individuals parameters to determine the effective dosingregimen of FGF-1 to be used in clinical trials. The epithelializationrate, which approximates the overall clinical healing rate, was giventhe highest priority. Since the amount of angiogenesis (Bv, NB) and thefibroblastic response (Fv) play an important role in the quality ofhealing, these parameters were next in priority. The time course of theinflammatory response (Mv and Nv) is also important in the determinationof the optimal dose of FGF-1.

Epithelialization rate was significantly enhanced by the growth factortreatments; indicating a faster closure of the wound by growth factortreatments at the four day period. This enhanced epithelialization rateresponse illustrates the faster rate of closure in the wound by growthfactor treatments, at least in the early stages of healing (up to 30%re-epithelialization). Angiogenesis was markedly higher in the initialphase in the growth factor treated groups. This angiogenic response wasnot statistically different between the treatments and the control atday eight. This implies that FGF-1 is angiogenic and triggers theformation of blood vessels earlier in the treated wounds.

Increasing the rate of angiogenesis in the wound center implies that therate of wound healing will be enhanced. Once the vessels infiltrate thecenter of the wound, which is normally low in oxygen, residentfibroblasts will produce more collagen and the epidermal repair rateshould increase. Since blood vessels are responsible for transport ofnutrients, immunoglobulin and white blood cells in the wound bed,resistance to infection is also increased by the enhanced angiogenicresponse.

A higher Fv in the 8 mg dose was seen at the wound center after eightdays; although the same dose showed a significantly lower Fv than theother groups at day four. The higher Fv at eight days is likely toindicate an increase in the amount and rate of collagen productionleading to an increase in overall wound strength. Also a good indicationof the healing response is the inflammatory response (Nv & Mv). Aninitial high response in the growth factor treatments compared to thecontrol indicated an initiation of an early inflammatory phase comparedto the control ulcers.

In general, the growth factor treatments affected the healing responseexhibiting a dose dependent behavior The addition of FGF-1 led to anincrease in the angiogenic and fibroblastic response as well as anincrease in the epithelialization rate. Based on the prioritizedparameters, the preferred dose was 8 mg. This dose initiated a highepithelialization rate, fibroblastic and angiogenic response and was thelowest dose to initiate these responses. In order to compare the resultsto the topical regimes used in previous studies, it is necessary todetermine the release kinetics of this system. Since FGF-1 has a shortin vivo half life, the use of the scaffold serves to protect it until itis released. Therefore, this fibrin matrix serves as a controlledrelease drug delivery system, while providing an adherent tissuescaffold during healing.

EXAMPLE 13 Delivery of Rate of FGF-1 Through a Porous Fibrin Scaffold ina Full Thickness Defect in a Rabbit Model

Controlled delivery refers specifically to the precise control of therate by which a particular drug is released from a system without theneed of frequent or repeated administration. Drug release rate, that isconstant, over a fixed prolonged period of time follows zero orderkinetics in which the rate is not affected by concentration.Biodegradable polymer matrices offer a major advantage in their use inthe improvement in the design of controlled release devices forimplantation as they may degrade in vitro to non-toxic products whichare readily eliminated in the body. The advantage of controlled deliveryin wound healing applications is to reduce problems in patientcompliance and reduce costs of wound management while enhancing the rateof healing by delivery of bioactive factors. Controlled delivery alsoreduces the exposure of the patient to excessive dose of the drugrequired at the targeted site.

Acidic fibroblast growth factor has been demonstrated to modulate thehealing response. A dose dependent behavior with the concentration ofthe growth factor has been observed. The different dose responsesindicate that there is a release of FGF-1 in the wound from the fibrinscaffold. A dose of 8 mg/ml enhances an angiogenic and fibroblasticresponse. This dose was preferred because it was the lowest dose whichinitiated the best response. FGF-1 is released when fibrin isenzymatically degraded.

EXAMPLE 14 Lodination of FGF-1

Pooled donor fibrinogen and bovine thrombin were prepared as describedabove. Polyethylene glycol beads and FGF-1 were prepared as describedabove. FGF-1 was iodinated by the IODO-GEN method as described byNeufeld with some minor modification (Neufeld et al., 1980). IODO-GEN(Pierce, Rockford, Ill.) was dissolved in CH₂Cl₂ to a concentration of 5mg/15 ml. 5 mg/15 ml of this solution was added to a 1.5 mlpolypropylene tube and evaporated to dryness under a stream of nitrogenwith FGF-1 (10 mg in 4.13 ml of 20 mM Tris-HCl, pH 7.5, 1 mM EDTA and1.5 M NaCl) together with 100 ml of 0.2 M Sodium Phosphate (pH 7.6). Thereaction was started by the addition of 1 mCi (1 Ci=37 GBq) IODO-GEN.After ten minutes at room temperature, the reaction was stopped by theaddition of 60 ml of 0.1% sodium metabisulfite and 30 ml of 0.1 mM KCl.The reaction mixture was then poured on to a Sephadex G-25 column (PD-10Pharmacia, Piscataway, NJ 08855-1327) for separation (Rizzino et al.,1988). The reaction tube rinsed with a equilibration buffer (10 mMHEPES, from a 1.5 M stock solution, pH 7.4), 0.3 M NaCl, 0.1% BSA andthe rinse and was again loaded onto a Sephadex G-25 column. The labeledFGF-1 was eluted from the column with 10 mM HEPES buffer (pH 7.4, 0.3 MNaCl). The radiolabeled I-125-FGF-1 peak (from the Sephadex G-25 column)was then loaded on the heparin-sepharose CL-6B column (Pharmacia LKB,Uppsala, Sweden) for further purification. The labeled I-251-FGF-1 wasrecycled ten times over the column to ensure maximal binding. Theheparin-Sepharose column was then washed with approximately 20 ml of 20mM Tris HCl (pH 7.5, 1 mM EDTA (TE)) buffer, until low counts elutedfrom the column. The column was then washed with a 0.5M NaCl buffer inTE. I-125-FGF-1 was eluted from the column with elution buffer 2M NaClin TE. The purification was affinity chromatography was 31.5%. Thespecific activity of the I125-FGF-1 was 31.5 mCi/mg. 350 mg FGF-1 wasadded to the I-125-FGF-1 solution and was mixed together for animalexperimentation.

EXAMPLE 15 Surgery

The rabbits were shaved and a depilatory agent was applied a day beforesurgery to ensure adhesion of the dressings. Four 3×3 cm full-thicknessdefects down to the panniculus camosus muscle were created on the dorsumof each of five white New Zealand rabbits. There were five periods ofstudy (1,2,5,8 and 16 days) with one rabbit per time period. Thisexperimental design gave a n value of four per time period. Thefull-thickness defects on the dorsum of the rabbit were covered with thefibrin scaffold containing I-125 labeled FGF-1. The growth factor wasadded to fibrinogen prior to mixing, at the concentration of 8 mg/ml(Smith & Nephew, Inc., Massillon, Ohio) wound dressing. A transparentpolyurethane, water vapor permeable, oxygen permeable, dressing (Opsite)was placed on these scaffolds. Dressings were secured with surgical tapearound, the perimeter. A cotton orthopedic stockinette, cut to fit, wasplaced around the animal along with a torso sling suit (Alice KingChatham, Hawthorne, Calif.).

EXAMPLE 16 Necropsy

At the time of necropsy, the animals were euthanized. The fibrin film,the dressing and the wound site were retrieved at the pre-determinedtime period to assess release kinetics. Radioactivity was measured usinga Picker Spectroscaler IIIA Gamma Counter (Picker X-ray Corporation,Cleveland, Ohio). Three readings of each of the four samples per timeperiod were averaged. This radioactivity in the fibrin scaffold was usedto determine the release rate of FGF-1. The radioactivity in the woundtissue and the wound dressing were also used to assess the growth factordistribution over time. A control in vitro reading of FGF-1 in a fibrinscaffold was taken to determine the initial amount of radioactivitypresent. A control wound at time zero was also made during necropsy andthe fibrin scaffold containing FGF-1 was placed on the wound. In thiscase, after in situ polymerization the fibrin scaffold was immediatelyretrieved to determine the amount of FGF-1 present in the fibrin clotafter implantation.

A wound model was selected by size and depth to take at least threeweeks to heal. Full thickness skin defects in the rabbit model were usedto simulate the initial phase of healing of skin wounds such as pressureulcers and burns. A wound down to the level of the panniculus carnosus,was chosen to simulate human skin; since skin mobility is reduced due tothe attachment to the muscle layer.

It was anticipated that the release rate of FGF-1 would be tied to thedegradation of fibrin. Therefore, the amount of FGF-1 added was equal tothe amount desired to be released over a one week interval.Incorporating FGF-1 in the fibrin scaffold during the clotting processis similar to intrafibrillar entrapment. It was also postulated thatthis release would occur only when cellular infilteration had takenplace thus making it a biofeedback mechanism. Since, FGF-1 has a shortin vivo half life, the use of a fibrin scaffold serves to protect isuntil it is released, thereby serving as a controlled release drugdelivery system, while providing an adherent tissue scaffold duringhealing.

An immediate release of about 35% of the total amount of FGF-1 was seen.This loss could be attributed to immediate absorption of FGF-1 during insitu polymerization of fibrin. Hence, all of the release cannot beattributed to degradation of fibrin. On day one there was a release ofup to 65% of the initial amount. This release could be accounted todiffusion of unbound FGF-1, and partly by enzymatic degradation of theporous fibrin scaffold by invading inflammatory cells. Thereafter, asustained controlled release was seen in the subsequent time periods.This sustained release can only be attributed to enzymatic degradationof the scaffold. Degradation of the scaffold also proceeds in acontrolled manner because of the type of cells present in the differentphases of healing. It appears that this scaffold degrades completelybetween day eight and day sixteen approximating the time for thesewounds to heal in this system. Hence, the initial release is diffusioncontrolled and the later release (after day one) is degradationcontrolled. The amount of FGF-1 released per day showed similar trends.An initial high release followed by constant release was seen. Thus, theporous fibrin scaffold served as an excellent vehicle for this growthfactor.

EXAMPLE 17 Use of Acidic Fibroblast Growth Factor (FGF-1) DeliveredThrough a Porous Fibrin Scaffold to Enhance Meshed Skin Graft Healing

Surgeons attempt to re-establish homeostasis in burn patients throughmeshed autologous skin grafts. Because these grafts contain noanatomical continuity with the donor site, the take of these graftsdepends on rapid establishment of a functional circulation and adherenceto the wound bed. Fibrin, the native adhesive in wounds, also acts as ascaffold for subsequent repair processes. Porous fibrin matrices havepreviously been shown to increase both the angiogenic and fibroblasticresponse. Additionally, fibrin matrices have been used as drug deliverydevices, acidic fibroblast growth factor (FGF-1) has been incorporatedinto this degradable fibrin scaffold.

For this study, an initial fibrinogen concentration of 60 mg/mL wasadded to thrombin in a 1:1 ratio giving an ultimate concentration of 38mg/ml. Water-soluble poly(ethylene oxide) (PEO) beads (17-24 microns indiameter, 2×10⁶ MW) were added to the clot at 25% by volume to make theporous matrices. FGF-1 was used at a concentration of 10 mg/mL.

Five 0.75×0.75 inch full-thickness defects were created on the dorsum ofeach of twelve white New Zealand rabbits: five rabbits in each of twotime periods (three and ten days). Skin harvested by dermatome wasmeshed (3:1) and cut to fit with the full-thickness wounds and attachedvia a fibrin scaffold, a porous fibrin scaffold, a porous fibrin/FGF-1scaffold, or 4-0 prolene sutures. Fibrin was used at 39.5 mg/ml.Fibrin/FGF-1 was used at 38.5 mg/ml and 10 micrograms/ml, respectively.The concentration for porous fibrin was 35.8 mg/ml while theconcentration of porous fibrin/FGF-1 was 34.7 mg/ml and 9 micrograms/ml,respectively. The wounds were covered with Tegaderm and the animals weresacrificed at three and ten days.

Upon sacrifice, the volume percent of cell types and tissue compositionwas determined. The healing rate, or the distance the wound marginmoves, was calculated and is representative of the amount of newgranulation tissue formed. For grafts, the muscle was excised to assayfor newly formed tissue at a strain rate of 0.2 inches per minute. Thefibrin tissue adhesive testing included porcine skin (0.01 inchesthick). The volume of glue/area skin was similar to that used in vivo.The strain rate used was 2.0 inches per minute.

Over half of the porous fibrin grafts did not adhere. The sutured graftsshowed a stiffer modular when compared with other treatments. Thenon-porous fibrin glue showed higher strength that the porous fibringlue at five and forty-five minutes. The fibrin/FGF-1 treatment resultedin a low modulus at three days.

The porous fibrin matrix served not only to adhere the graft better butalso to speed the healing rate versus the control. A relatively constantrate of FGF-1 release was observed and an increased angiogenic responsewas found when compared to the control.

EXAMPLE 18 Comparison of Shear Strength of Fibrin and Albumin Glues

The shear strengths of human abdominal skin adhered by either fibrin(1:1 ratio of fibrinogen to thrombin) or albumin, stabilized bycross-linking with polyethylene glycol (PEG), (25% albumin-5% PEG or 25%albumin-20% PEG) glue were compared. The fibrin had a shear strength of6.2 kPa, the 25-5 albumin about 9 kPa (about 1.5 times the fibrinvalue), and the 25-20 albumin glue appears to have a shear strength ofabout 31 kPa (about 5 times the fibrin shear strength). Thus, albuminglue appears to have a substantially higher shear strength than fibrinwhen used as an adhesive for skin grafts.

EXAMPLE 19 Comparison of Albumin and Fibrin Glues f6r Use With MetallicImplants

Bone loading stimulates the rate of bone integration at optimal levelsbut can compromise bone stability, at higher loads. Dental implants aretypically left unloaded for up to six months to minimize micromotion andto allow for bone integration and long term success. In addition,fixation of a percutaneous device is an important factor in preventingnon-epidermal as well as epidermal failure. Fixation of dental implantshelps prevent failure; allowing a tight seal at the gingival-implantinterface minimizes micromotion as well as the risk of infection causedby breakdown at the implant surface.

The potential of using bioadhesives to reduce the time when an implantis unloaded, without compromising implant stability was examined. Thesebioadhesives also have the potential to serve as degradable scaffolds tofurther enhance tissue integration.

The adhesive shear strength of these bioadhesives is critical for bothload transfer and implant stability. As an initial investigation, theadhesive shear strengths of fibrin and albumin glues were tested forboth metal-metal and skin-skin cases.

Fibrin glue was made from a 1:1 ratio of fibrinogen(7000-9000 mg/dl) tothrombin (500 units/ml), buffered in 40 mM CaCl₂. Albumin glue consistedof 25% albumin—20% polyethylene glycol (PEG) or 25% albumin—5% PEG.Curing time after application was one hour for all systems.

For the skin-skin tests, the epidermal surface of human skin was adheredto aluminum jigs, using cyanoacrylate glue. The inner faces of the skinwere adhered to each other with fibrin, 25-20 albumin, or 25-5 albumin.For the metal-metal tests, the inner faces of the aluminum jigs wereadhered to each other with either fibrin or 25-20 albumin. The effectivearea of each face of the jig was 3 cm². The specimens were tested inaxial shear, at a strain rate of 2.5 cm/min, to determine the adhesiveshear strength.

For skin-skin specimens, the 25% albumin−20% PEG system had the highestshear strength value (about 5 times that of the fibrin glue system)while the 25-5 albumin system was 1.5 times as strong as the fibrinsystem (˜6 kPa). For metal-metal specimens, the albumin glue (˜35 kPa)was substantially stronger than fibrin (˜1.5kPa) p<0.004). Therefore,the albumin glue provided greater adhesive strength to both aluminum andskin. This difference was accentuated in the metal-metal case. For thealbumin glue, increasing the amount of PEG, used to crosslink the gel,significantly increased the adhesive strength. The use of albumin glue,therefore, may be a viable option for improving the stabilization andlong-term success of dental implants.

EXAMPLE 20 Effect of Composition of an Adhesive Albumin Glue onBiocompatibility and Healing in an Incision Wound Model in RabbitsMETHODS AND MATERIALS

Materials

Albumin solutions were prepared by dissolving rabbit albumin (Sigma, StLouis, Mo.) in a 0.85% sodium chloride solution. PEG-SPA solution wasprepared by solubilizing PEG-SPA (MW=20,000) in a basic HEPES solution.Specimens were prepared by combining 25 μL of each component. Fibrin wasprepared by combining 25 μL each of human fibrinogen (Sigma, St Louis,Mo.) dissolved in distilled water, and bovine thrombin (Sigma, St Louis,Mo.) solubilized in a 40 mM calcium chloride solution. Wounds weretreated as follows (2 wounds/treatment/animal): sutures (2/wound;control), fibrin glue, adhesive albumin (25% albumin solution, 15%modified PEG solution) (i.e., 25/15), adhesive albumin (25/10), andadhesive albumin (30/10).

Surgical Procedure

Eight New Zealand White rabbits were sedated with Ketamine and Xylazine,shaved, and cleaned with Betadine to prepare a sterile field. Rabbitswere anesthetized using Isofluorane. Five full-thickness incisionalwounds, 2.5 cm long, were created on each side of the spine. Theepidermis and dermis were cut, with care taken to avoid cutting theunderlying panniculous carnosus muscle. After blotting each wound toremove any blood, adhesive was applied (25 μL/solution). After glue wasallowed to clot in wounds for 15 seconds, excess was squeezed out andeach wound was held closed for one minute.

Evaluations

Wounds were visually examined daily and photographed every three or fourdays. Wound closure, defined as apposed skin with only a surface scab orno scab, was noted. At each of seven and fourteen days post-operatively,four rabbits were sacrificed. In four rabbits, all wounds were used forhistology. From two of the seven-day rabbits, and two of thefourteen-day rabbits, only five of the ten wounds were used forhistology, with the other wounds used for different analyses.Immediately post-sacrifice, strips approximately 6 mm×1 cm were excisedfor histological analysis and placed in 10% neutral buffered formalin.

Specimens were processed, embedded in paraffin, sectioned and placed onslides. To microscopically examine cellular response and the woundregion, specimens were stained with hematoxylin and eosin. Digitalimages (4×, 20× and 40×magnifications) were obtained using a microscope(A×70, Olympus, Melville, N.Y.) and digital camera (OLY-750, Olympus,Melville, N.Y.). From the 4×images, the wound region, materialremaining, and new tissue were traced using image analysis software(ImageProc Toolkit, Reindeer Productions, Raleigh, N.C.; AdobePhotoshop, Adobe Systems, Mountain View, Calif.). Material remaining andtissue infill were represented as volume fractions of the wound area,calculated by dividing the area of material remaining (or new tissue) byarea of the wound region. Wound region was defined as the areaoriginally created by the incisions (i.e., empty of tissue at timezero). Wound closure, as indicated by the presence of a continuousepidermis, was also noted. From the 20×images, volume fraction of bloodvessels was measured. Percent of neutrophils, macrophages andfibroblasts (i.e., as a percent of the total of all three) wasdetermined from the 40×images.

RESULTS

Wound Closure

Results from both the visual and microscopic analyses at seven andfourteen days were similar. In one rabbit, seven of the ten wounds didnot close. By day one, three of the wounds appeared infected and werescrubbed and received a topical antibiotic. No wounds on this rabbitwere closed at one week. After eliminating this rabbit, all wounds wereclosed by one week—as determined from histological specimens—except fortwo sutured wounds and two fibrin treated wounds. All wounds, excised attwo weeks had closed. The resultant percent of closed wounds pertreatments, as seen in histological specimens, shown in Table 1.

TABLE 1 material closed remaining tissue infill epidermis (%) (%) (%wounds) treatment 1 week 2 weeks 1 week 2 weeks 1 week 2 weeks suturesn/a n/a 87.3 ± 91.3 ± 66.7 100  9.4 5.2 fibrin 3.65 ± 0.04 ± 78.1 ± 95.2± 66.7 100 6.32 0.11  7.6 3.8 25/15 10.4 ±  2.14 ± 73.1 ± 89.7 ± 100 1004.49 5.23 10.0 11.9  25/10 7.01 ± 1.83 ± 79.9 ± 92.6 ± 100 100 7.94 2.2310.3 6.1 30/10 4.48 ± 0.00 84.0 ± 96.7 ± 100 100 4.39 14.5 2.8

Material Remaining

Fibrin and 30/10 adhesive albumin degraded fastest, at similar rates,with a mean of 3.6% and 3.4% material-remaining at one week,respectively. By two weeks; there was less than 0.5% of 30/10 albumin orfibrin remaining. Increasing PEG-crosslinker concentration sloweddegradation at one week, with 12.3% material remaining in the 25/15treated wound vs. 7.0% remaining in the 25/10 treated wound. At twoweeks, a mean of 2.1% and 1.8% material remained in the 25/15/ and 25/10closed wounds, respectively, as shown in FIG. 2.

Tissue Infill

In some specimens, there was a region devoid of any evident tissueingrowth or material. In most cases, the open areas were surrounded by aslightly denser layer of cells, suggesting a capsule formation. Thisregion may have been filled with fluid, or with tissue or material thatwas removed in the histological processing. For this reason, a tissueinfill is presented separately from material remaining.

Referring to FIG. 3, by one week, wounds in all treatment groups werefilled with at least 75% new tissue. The suture-closed wounds werefilled 96.4±1.5%, on average. The 25/15 albumin treated wounds had theleast infill at one week (mean 76.7±10.0%). Fibrin treated wounds showedan infill of 92.3±11.3% at one week, which was slightly less than thatseen with the 10% PEG-SPA albumin materials (25/10 and 30/10), althoughnone of the differences were significant. By two weeks, wounds were atleast 90% filled, on average, with the sutured wounds possessing theleast mean infill (89.7±9.1%). Wounds treated with the 25/15 albumindemonstrated a mean infill of 94.1±5.5%, similar to the fibrin treatedwounds (94.2±3.4%). The 30/10 albumin treated wounds showed the mostinfill at two weeks, with an average 96.7±2.8% infill.

Many wounds demonstrated a slight “depression” at the surface. There wastypically either albumin or scab material was present in this depressedarea. At one week, this area was largest (mean 14.2±11.1% wound area) infibrin treated wounds, and smallest in the 25/15 and 30/10 treatedwounds (1.9±1.6% and 2.1±3.6% wound area, respectively). At two weeks,this region was largest in the suture treated wounds (6.5±3.5%), asopposed to the 25/10 and 25/15 albumin closed wounds (1.7±1.4% and1.8±1.7% wound area, respectively).

Inflammatory Response

Inflammatory cells in the wound region included neutrophils andmacrophages. Other cells, including lymphocytes and giant cells, wereseen in very small quantities in some wounds as shown in FIGS. 4A and4B.

DISCUSSION

Wound Closure

By two weeks, all wounds had a complete epidermis, indicating that alltreatments, including the control (sutures), are effective closuretreatments. While epidermal closure is one measure of wound healing, itwas evident from the histology that epidermal closure did notnecessarily constitute complete material metabolism or tissue infillwithin the dermis.

Material Remaining

Use of a degradable material allows for replacement of the material bynew tissue, resulting in complete tissue infill, constituting completehealing. A degradable material should degrade slow enough to remainpresent long enough to provide adequate structural support (e.g., as amechanical structure or tissue scaffold), or to deliver a drug at anappropriate rate and duration, to facilitate optimal healing. It shouldalso degrade fast enough to not interfere with new tissue growth.

For the albumin materials, the goal was to achieve material degradationand resolution of the inflammatory response faster, or at least asquick, as with fibrin. In the wounds treated with the albumin adhesivewith the highest concentration (30%) of the natural component, albumin,the degradation rate was similar to that of fibrin. Slightly slowerdegradation was associated with the material with increased crosslinkercontent (i.e., 25/15 vs. 25/10).

Tissue Infill

Because it is not known whether or not the empty areas in some specimenswere filled with fluid, cells or remaining adhesive, it is not possibleto draw definitive conclusions about the subepidermal histology.

Inflammatory Response

While inflammatory response is required for normal wound healing, aprolonged inflammatory response may indicate a chronic pathologicalcondition. The inflammatory response must be resolved before completehealing can occur. The inflammatory response observed here in thealbumin treated wounds did not resolve as quickly as in the fibrintreated wounds. Based on rate of resolution between one and two weekshowever, it is expected that the inflammatory response would becompletely resolved by three weeks or before.

Blood Supply

In all wounds, blood supply, as measured by volume fraction of bloodvessels, decreased between one and two weeks. At one week, volumefraction of blood vessels was highest in the 25/15 treated wounds (mean50.9%) and lowest in the suture wounds (42.9%), with values in allalbumin treated wounds higher than both the sutured and fibrin treatedwounds. Between weeks one and two, mean blood vessel volume fractiondecreased between 31-41% in all treatment groups except the 25/10 group,which decreased only 20%.

CONCLUSIONS

Although tissue infill was not complete at one or two weeks, epidermalclosure occurred in all albumin treated wounds by one week, and in allwounds by two weeks. In both sutured and fibrin treated wounds, theepidermis of 33% of the wounds was not complete by one week. Tissueinfill in both, however, was slightly higher in those wounds at oneweek. By two weeks, tissue infill in all wounds was over 90%. Of thealbumin materials, highest infill was associated with a higher albuminconcentration, suggesting that higher content of the natural componentof the adhesive induced faster metabolization.

While some cells were present in the albumin adhesive, this occurrencewas not common. Degraded areas around the cells suggest that the cellswere metabolizing the material; In some instances, cells were present inmaterial on the outside of the wound.

Wounds treated with albumin were more likely to have a smallerdepression at the external surface of the wound, suggesting a smootherscar. The largest mean area of depression was associated with fibrin(14% wound area) at one week, and fibrin at two weeks (6% wound area).

EXAMPLE 21 Effect of Porosity on Shear Strength of an Albumin AdhesiveMETHODS AND MATERIALS

Preparation of Materials

Adhesive albumin was prepared by combining an albumin solution andsolubilized crosslinker polymer, 1:1 (v.v). Albumin solution wasprepared by dissolving human albumin (Sigma, St Louis, Mo.) in a 0.85%sodium chloride solution. PEG-SPA crosslinker solution was prepared bysolubilizing PEG-SPA (MW=20,000) (Shearwater Polymers, Huntsville, Ala.)in a basic HEPES solution. Specimens were prepared by combining the twofluid components 1:1 (v/v).

For porous specimens, polyethylene glycol (PEG) particles (MW=100,000)(Sigma Chemical, St Louis, Mo.) were sprinkled on the adhesive layerafter mixing the albumin and modified-PEG solutions. The polymerparticles dissolve immediately, leaving a void. Particles were dividedby size by sifting PEG particles through nylon meshes of known spacing.The “small” particles were separated out using 40 μm and 120 μm meshes,and the “large” particles sorted using. 120 μm and 200 μm meshes. Toestimate particle volume of each pore size required for 10% and 20%porosity, densities of the two size ranges of particles were calculatedby weighing 100 μL particles.

Examination of Structure

Both the particles and the structure of adhesive albumin clots obtainedwith different particle sizes were examined using light microscopy. Forexamination of the particles, particles were randomly sprinkled onto aglass slide. One to three images of the particles were digitized at10×magnification (Provis A×70, Olympus, Melville, N.Y.) to provide for aminimum of 30 and 50 samples of small and large particles, respectively.Particles were traced using image analysis software (Adobe Photoshop,Adobe Systems, Mountain View, Calif.; ImageProc Toolkit, ReindeerProductions, Raleigh, N.C.) to obtain particle area. Approximatediameter was calculated from the area value of each particle. A meandiameter value was calculated for the small and large particles.

Clot specimens (30 μL spread over 1 cm²) with small or large pores werecreated on glass slides, and examined at 10×magnification. Again, imageswere digitized and pores traced using image analysis software. Each porewas traced, area measured, and values calculated, as for the particles.Three sample groups of small pores and five sample groups of largerpores were traced to provide a minimum of 50 and 30 samples,respectively.

Evaluation of Shear Strength

Porcine dermis (Brennen Medical Inc, St Paul, Minn.) (6.25 cm².) wasglued, hair side down, to aluminum jigs using cyanoacrylate glue asshown in FIG. 5. The glued skin pieces, kept moist by a covering ofsaline-soaked gauze, were allowed to set for fifteen-thirty minutes.After the moist gauze was removed and the skin pieces were blotted withdry gauze, albumin solution and modified-PEG solution were applied toone piece of skin and mixed with one applicator tip. Adhesive albuminwith four different porosity configurations was tested: no pores; 10%volume porosity, small pores (10/small); 20% volume porosity, smallpores (20/small); and 20% volume porosity, larger pores (20/large). Foreach specimen, a total of 188 μL volume of material was applied betweenthe pieces of skin, for an approximately 300 μm layer of adhesive.

Following application of each test material, specimen pairs were pressedtogether and allowed to set for one hour, with a 240 g weight appliedacross five specimens to allow for even setting. After setting,specimens were clamped (Quick-Grip Micro Bar Clamp and Spreader,American Tool Companies, USA) to prevent specimen damage during loading,and loaded into custom prepared fixtures in an MTS Minn Bionix 858System (MTS Systems, Eden Prairie, Minn.) After clamp removal, specimenswere tested in shear at 2.5 cm/minute until failure. Load anddeformation were recorded. Three runs were performed, with fivespecimens per sample type tested in each run.

Force-deformation plots and stress-strain plots were generated and areshown in FIGS. 6A-B. Values were calculated as follows:

τ_(ult)=ultimate shear strength=F_(max) /A

γ_(xy)=shear strain=tan (dy/x)≈dy/x

G=shear modulus=(τ_(0.8 ult)−τ_(0.2 ult))/(γ_(xy 0.8)−γ_(xy 0.2))

W=failure energy=∫F dy

Where F_(max) = maximum attained load A = bonded area of skin = 6.25 cm²dy = deformation (in direction of load) x = material thickness ≈ volumeadhesive applied/A = 300 μm

Standard statistical values, including mean, standard deviation andstandard mean error (±1) were calculated. ANOVA analyses with Fisher'sprotected least significant difference test were performed (p<0.05). Todetermine the failure mode of the adhesive materials, specimens werevisually examined after testing to identify distribution of thematerials. Because it was suspected that humidity and temperature mayaffect material shear strength, each was recorded each day of testing.

RESULTS

Examination of Structure

For both the small and large particles, mean particle size wasapproximately equal to the theoretical average particle size, calculatedas the median of the particle size range. Theoretical median diameter ofthe small (40-120 μm diameter, FIG. 7A) and large particles (120-200 μmdiameter, FIG. 7B) was 80 μm and 160 μm, respectively. Mean of the,small particle samples (n=51) was 79±16 μm. Mean of the large particles(n=31) was 160±40 μm.

In specimens prepared with small particles, mean pore size was 97±28 μmdiameter (n=62). In specimens prepared with large particles, pores were193±57 μm diameter (n=37). The range of pore sizes was 48-181 μm, and105-363 μm, in the small and large pore clots, respectively. Resultantmean pore size, as observed with light microscopy, was approximately 23%and 21% larger than mean particle size, for small and large particles,respectively. See FIGS. 8A and 8B.

Evaluation of Shear Strength

Data from individual runs was pooled for analysis. Data from eachindividual run was also examined and it was confirmed that trends seenwithin each run were similar to those in the pooled data. Of the fifteenspecimens tested per sample type, between one and three specimen resultswere eliminated from analysis due to damage during testing. Specimenresults were eliminated for the following reasons: peeling of skin frommetal jig; visible motion in the glue joint during loading into thefixture; jamming of the test jig during loading; or, visible evidence ofaccumulation of adhesive or sticking of adhesive at the end(s) of aspecimen. Up to an additional two results from each sample group wereeliminated as outliers per the boxplot rule (Devore, Probability andStatistics, 4^(th) Ed., Boston, Duxbury Press). Stress-strain curveswere typically slightly asymmetric, with the ascending slope steeperthan the descending slope (see FIG. 9).

As expected, shear strength of the nonporous adhesive albumin(10.83±4.92 kPa) was stronger (28-35%) than the porous materials (Table2), although the difference between the non-porous and porous materialswas significant (p<0.05) only for the 90/small material. There were nosignificant differences in ultimate shear strength between the threeporous materials.

TABLE 2 τ_(ult) ultimate γ W G shear strain at failure shear strengthultimate energy modulus material n (kPa) strength (J) (kPa) no pores 1010.83 ± 4.92 3.47 ± 0.78 10.41 ± 4.72  9.57 ± 6.20 10/small 14  8.05 ±2.59 2.19 ± 0.66  6.04 ± 1.84 14.73 ± 4.56 20/small 13  8.40 ± 2.79 1.78± 0.67  5.89 ± 2.26 15.54 ± 4.31 20/large 10  8.17 ± 1.75 2.31 ± 0.75 5.77 ± 0.72 18.02 ± 7.24

Strain at ultimate strength was significantly higher for the nonporousadhesive (3.47±0.78 kPa) than for the porous materials (see FIG. 10).While these differences in tolerated shear strain of the porousmaterials were not significant, the 20/small pore material tolerated theleast (1.78±0.67 kPa).

With the testing method used, variation in the slope of thestress-strain curve, or the shear modulus, was large. Despite this,shear modulus of the material with no pores (9.57±6.20 kPa) wassignificantly lower than that of the porous materials (p<0.03).

Temperature and humidity varied only slightly on the test days, withtemperature and humidity values in the lab of 72-76° F. and 42-47%,respectively. Results varied slightly from run to run, but did notcorrelate with temperature or humidity differences.

Visual examination revealed that material remained on both pieces ofskin of each sample pair, and that, except for isolated patches in thematerial layer on a few specimens, remaining material covered a majorityof each skin piece. This distribution suggested cohesive, rather thanadhesive failure, since the failure appears to have occurred within thelayer of adhesive rather than at the interface between the glue andsubstrate.

DISCUSSION

In previous studies of porous vs. non-porous fibrin adhesive to secureskin grafts in rabbits, it was found that porosity was associated withincreased graft removal (Osborne et al. (1996) Trans. Soc. Biomatls.I:27). There was a 30% decrease in shear strength when pores were formedat a 12% volume, using particles 170-240 μm diameter. Pores in fibrinclots were associated with failure of 40% of grafts at both three andten days in grafted wounds in rabbits, although only 20% of porousfibrin grafts failed at ten days when acidic fibroblast growth factor(FGF-1) was included in the clots (Osborne et al. (1996) Trans. Soc.Biomatls. I:27).

This was a primary motivator for this study. By better understanding andanticipation of the reduction in strength associated with changes inporosity, materials may be developed to better preserve strength.Selection of PEG-derivative crosslinked adhesive albumins, which havebeen observed by some to possess superior adhesive compared to fibrin(Truong et al. (1996) Trans. Soc. Biomatls. II-73; Barrows et al. (1996)Trans. Soc. Biomatls. I-8; Huang et al. (1997) Trans. Wound Healing Soc.7:63), was intended to compensate for some of the reduction in strengthby providing a stronger material to start with. Even without inducedmacropores, fibrin has been found by some to have insufficient strengthfor some applications (Gibble et al. (1990) Tarnsfus. 30(8):741-47;Truong et al. (1996) Trans. Soc. Biomatls. II-73).

Loss of grafts may lead to additional surgery and prolongation ofhealing. Additional surgery implies added cost, extended hospital stay,and requirement of additional skin for grafting. If insufficient skin isavailable from the patient, this may require waiting for previously useddonor sites to heal. Prolonging time to coverage and healing of woundsmay also result in increased chance of infection.

As previously eluded to, while adequate adhesion is a primary concern,so too is sufficient penetrability of the material to ensure that cellscan move in and replace the material easily. While this can beaccomplished by inclusion of pores, this can also be achieved byproviding a material that degrades in synchronization with the rate ofnew tissue growth. In most degradable materials, this rate ofdegradation can be increased by increasing material surface area, suchas with pores. Thus, incorporation of pores in a degradable material notonly better facilitates immediate passage through the material, but alsocan be used to adjust the degradation rate.

Examination of Structure

Two different pore size ranges were selected, to determine the effect ofsize on material strength. To ensure tissue ingrowth into materialplaced in a wound, pores should be at least 40-50 μm in diameter(Feldman et al. (1983) Biomatls. 4(2):105-111; Morehead et al. (1994)Otolaryng. Clin. North Am. 195-201) and optimally 100 μm (Dagalakis etal. (1980) J. Biomed. Mater. Res. 14:511-528). Also, bacteria can growinto pores greater than 1 μm, but phagocytic cells needed to clear awaybacteria cannot penetrate pores smaller than 50 μm (Morehead et al.(1994) Otolaryng. Clin. North Am. 195-201). Therefore very small pores,such as those in unmodified fibrin (i.e., less than 10 μm) (Blomback etal. (1990) Fibrinogen, Thrombosis, Coagulation, and Fibrinolysis, NY:Plenum Press); do not allow ingrowth and are more likely to harborinfection (Morehead et al. (1994) Otolaryng. Clin. North Am. 195-201).

The ideal pore size in materials placed on wounds is not known.Theoretically, expanding pore size may increase tissue ingrowth bywidening the pathway for new tissue. On the other hand, this may alsoincrease degradation rate by expanding the amount of material surfacearea and may decrease shear strength by reducing contact area betweenthe material and substrate.

The particle sizes used were selected to produce pores approximatelybetween 50 and 200 μm, with median values equally spaced within thatrange, assuming that mean diameter would be similar to the median, andthat pore size would approximately correspond to particle size. Meansmall and large particle size, measured from light microscopy images,corresponded well with the median values of each particle range. Poresizes associated with each particle size range were approximately 25%larger than mean particle size, suggesting a slight expansion in thevoid associated with dissolution of the particle. Standard deviation ofpore size (29-30%) was only slightly larger than that of particle size,implying that the distribution is similar.

One disadvantage of the methods of pore examination used here is thelack of a cross-sectional view. Future studies may involve examinationof the matrices using confocal microscopy. Without this distribution, itis not possible to conclude that pores are, interconnecting and it islikely that they are not. Design of the material, however, is based onthe premise that the material is degradable. It is assumed that thininterstices between pores will degrade in a short amount of time,allowing for passage of cells, chemicals and new vascular andcollagenous tissue.

Evaluation of Shear Strength

As expected, inclusion of pores into the material reduced ultimate shearstrength (significance at p<0.08). Differences between the three porousmaterials were slight, however. Shear strain and failure energy of thenonporous material was significantly higher (p<0.001) than that of theporous materials, which reduced shear strain by 37, 49 and 33%(10/small, 20/small and 20/large, respectively). These results suggestthat shear strain of a 20% porous material can be improved either byincreasing pore size or decreasing pore volume.

Shear modulus (i.e., slope of the elastic region of the stress-straincurve) of the nonporous material was significantly lower (p<0.03) thanthat of the porous materials (35-47% lower). This difference may be, inpart, due to the effective reduction in contact area between thematerial component of the matrix and the substrate, with the inclusionof pores. The area of the pores at the substrate was not considered incalculation of the strength or other values since our interest was inthe properties of the resultant scaffolds. Again, differences betweenthe porous materials were only slight.

Differences among the porous matrices were most pronounced for shearstrain and shear modulus. Material with 20% porosity/small pores (i.e.,20/small) showed lower shear strain than both the lower porosity(10/small) and the larger pore (20/large) materials. These resultssuggest that shear strain can be improved either by increasing pore sizeor by decreasing pore volume. A high shear strain is desired because itreflects how much the material can deform before it breaks. While highshear strength is required to protect material from failure with highshear forces, high shear strain at ultimate strength can aid inresistance to failure from small shear forces acting on the grafts, suchas with slight patient movement or healthcare worker handling.

There are theoretical structural implications associated with changingeither porosity or pore size. Using the mean diameters for small andlarge pores measured from light microscopy images, and approximatingeach pore as a sphere, surface area of the pores within a material canbe calculated, for a given pore volume (V_(matrix)) (e.g., 188 μL).Surface area (SA) can be calculated as SA=N×SA_(single pore), whereN=number of pores=V_(all pores)/V_(single pore). For this study,V_(all pores) is equal to either 0.1V_(matrix) or 0.2V_(matrix) (10 or20% porosity, respectively). This increase in surface area translates toan increase in length and/or decrease in thickness of interstices.

Assuming the same volume of porosity (e.g., 20%), theoretical totalsurface area (SA) in a matrix with small pores is approximately 1.97times as great as that with large pores, (i.e.,SA_(small):SA_(large)=1.97:1). Assuming one uniform pore size, as thevolume of porosity is increased from 10% to 20%, the correspondingtheoretical surface area of the pores is doubled (i.e. SA₂₀:SA₁₀=2:1).

These theoretical surface area differences, which imply a change inlength of interstices, do not directly correlate with results obtainedhere. Expansion of the interstices between pores implies that given somedegree of material elasticity—the matrix should tolerate moredeformation. As previously mentioned, however, this increase in voidvolume and area results in reduced material (i.e., as opposed to voidspace) at the surfaces subject to the load. In the porous scaffolds,larger pores resulted in higher shear strain (i.e., 20/large vs.20/small). This shear strain was similar to that of the 10/smallscaffold. The fact that the highest shear strain was obtained with thenonporous material indicates that inclusion of the pores weakened thescaffold too much to overcome failure with increased deformability.

The PEG-SPA crosslinked albumin adhesive used here was a slightlydifferent composition than that previously studied by Huang andassociates (Huang et al. (1997) Trans. Wound Healing Soc. 7:63) inearlier shear strength tests. With a 25% albumin-20% PEG-SPA (i.e.,25/20) system on skin, shear strength was approximately five timesstronger than that of the fibrin glue (75-90 mg/mL fibrinogen, 500 U/mLthrombin). Shear strength of the 25/5 system on skin was approximately1.5 times stronger than that of the fibrin system. The compositionselected for use here was based on degradation and biocompatibility, inaddition to strength, as seen in a study of the material in incisionalwounds in rabbits. The subsequent selection of a composition of 30%albumin/10% PEG-SPA was based on degradation properties and resolutionof inflammatory response. All materials studied resulted in woundclosure. The 30/10 material possesses slightly lower shear strength thanthe 25/15 system, for example, as well as compared to the 25/20 systemused by Huang and coworkers. On the other hand, it exhibited highershear strain at ultimate strength and approximately equal failureenergy, compared to the 25/15 material, suggesting similar toughness.This material was selected for its degradability rate similar to fibrin,and superior shear strain at ultimate strength. Although ultimatestrength is lower than for the 25/15 material, higher shear strain, withsimilar failure energy, suggests that the material tolerates more motionbefore failing.

CONCLUSIONS

In a PEG-SPA crosslinked albumin adhesive, with particles added tocreate pores, shear properties can be modified by changing particle sizeand particle volume added. Mean pore diameter was approximately 25%greater than the diameter of particles used to create the pores.Incorporation of three different configurations of pores reduced theshear strength by approximately 25% (i.e., 22-26%). Failure energy wasreduced by approximately 45% (i.e., 42-45%). There were no significantdifferences in any values for the porous materials. Shear strain wasreduced slightly both by increasing porosity with a given pore size, orby decreasing pore size for a given pore volume. Increasing pore sizeresulted in a slight increase in shear modulus, for a given volume. Ifinitial strength can be improved, this system may show promise forapplication as a tissue adhesive to stabilize skin grafts.

Throughout this application, various publications are referenced bycitation and number. Full citations for the publication are listedbelow. The disclosure of these publications in their entireties arehereby incorporated by reference into this application in order to morefully describe the state of the art to which this invention pertains.

The invention has been described in an illustrative manner, and it is tobe understood the terminology used is intended to be in the nature ofthe description rather than of limitation.

Obviously, many modifications and variations of the present inventionare possible in light of the above teachings. Therefore, it is to beunderstood that within the scope of the appended claims, the inventionmay be practiced otherwise than as specifically described.

In view of the teaching presented herein, other modifications andvariations of the present inventions will be readily apparent to thoseof skill in the art. The foregoing drawings, discussion, and descriptionare illustrative of some embodiments of the present invention, but arenot meant to be limitations on the practice thereof. It is the followingclaims, including all equivalents, which define the scope of theinvention.

REFERENCES

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BARROWS et al. (1996) Evaluation of a new tissue sealant material: Serumalbumin crosslinked in vivo with polyethylene glycol. Trans. Soc.Biomatls. I-8.

BLOMBACK et al. (1990) Native fibrin gel networks and factorsinfluencing their formation in health and disease. (C. Y. Liu and S.Chen, eds), Fibrinogen, Thrombosis, Coagulation, and Fibrinolysis, NewYork: Plenum Press, pp. 1-23.

BLUM et al. Effect of Composition of an Adhesive Albumin Glue onBiocompatibility and Healing in an Incisional Wound Model in Rabbits.

BLUM et al. Effect of Porosity on Shear Strength of an Albumin Adhesive.

CRANIN et al. (1996) 9.4. Dental Implantation. In: Biomaterials Science(B. D. Ratner, A. S. Hoffman, F. J. Schoen, J. E. Lemons, eds), SanDiego: Academic Press, pp. 426-435.

DAGALAKIS, et al. (1980) Design of an artificial skin. III. Control ofpore structure. J. Biomed. Mater. Res. 14:511-528.

DEVORE J. L. Probability and Statistics, 4^(th) Edition. Boston: DuxburyPress;

FELDMAN et al. (1983) Electron microscopic investigation of soft tissueingrowth into Dacron velour with dogs. Biomatls 4(2):105-111.

FELDMAN and von RECUM (1985) Non-epidermally induced failure modes ofpercutaneous devices. Biomaterials, 6(5)352-356.

GIBBLE et al. (1990) Fibrin glue: the perfect operative sealant?Transfus. 30(8): 74147.

HUANG et al. (1997) A comparison of the shear strength of fibrin andalbumin glues. Trans. Wound Healing Soc. 7:63.

KILPADI et al. (1998) A comparison of the adhesive strength of albuminand fibrin glues for use with metallic implants. Proc. of the 24thAnnual Meeting of the Society for Biomaterials, San Diego; Calif.

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What is claimed is:
 1. A bioadhesive scaffold material, said material comprising: albumin; a polyethylene glycol crosslinking agent; 10 to 50 volume percent of 10 to 500 micron polymeric beads imparting porosity, said beads selected from the group consisting of: polyethylene glycol and poly(ethylene oxide); and a modifying agent incorporated into said bioadhesive material, wherein said modifying agent is selected from the group consisting of: a tissue growth promoting agent fibroblast growth factor I (FGF-I), an anti-inflammatory and an antibiotic.
 2. A bioadhesive scaffold material according to claim 1, wherein said polyethylene glycol comprises a modified polyethylene glycol.
 3. A bioadhesive scaffold material according to claim 1, wherein said modifying agent incorporated into said bioadhesive material is said tissue growth promoting agent.
 4. A bioadhesive scaffold material according to claim 3, wherein said tissue growth promoting agent is said fibroblast growth factor I, (FGF-I).
 5. A bioadhesive scaffold material according to claim 1, wherein said modifying agent is said anti-inflammatory.
 6. A bioadhesive scaffold material according to claim 1, wherein said modifying agent is said antibiotic.
 7. A bioadhesive scaffold material according to claim 1, wherein the concentration of said albumin in said bioadhesive scaffold material ranges from approximately 0.1% to 50%.
 8. A bioadhesive scaffold material according to claim 7, wherein said amount of said albumin in said bioadhesive scaffold material ranges from approximately 1.0% to 30% by weight.
 9. A bioadhesive scaffold material according to claim 7, wherein said concentration of said albumin in said bioadhesive scaffold material ranges from approximately 10% to 30%.
 10. A bioadhesive scaffold material according to claim 1, wherein said concentration of said polyethylene glycol crosslinking agent in said bioadhesive scaffold material ranges from approximately 0.1% to 30%.
 11. A bioadhesive scaffold material according to claim 10, wherein said concentration of said polyethylene glycol crosslinking agent in said bioadhesive scaffold material ranges from approximately 5.0% to 25%.
 12. A method for adhering tissue comprising the steps of applying and bonding bioadhesive scaffold material of claim 1 to the tissue.
 13. A bioadhesive scaffold material according to claim 1 wherein said porosity imparting polymeric beads comprise polyethylene glycol.
 14. A bioadhesive scaffold material according to claim 1 wherein said porosity imparting polymeric beads comprise poly(ethylene oxide).
 15. A bioadhesive scaffold material according to claim 1 wherein said porosity imputing polymeric beads have a mean size ranging from 10 microns to 500 microns.
 16. A bioadhesive scaffold material according to claim 15 wherein said porosity imparting polymeric beads have a mean size of from 120 microns to 200 microns. 